A. Field of Invention
This invention generally relates to detectors for positron emission tomography (PET), and/or single photon emission computed tomography (SPECT).
B. Description of the Related Art
PET and SPECT imaging devices operate by sensing gamma radiation emitted by radiopharmaceuticals that have accumulated in target organs or tissues of a patient. A two or three dimensional image is constructed by mapping the positions of particular gamma sources. With specific reference to a PET experiment, a selected radiopharmaceutical is administered to a patient, which can comprise any of a wide variety of physiologically relevant molecules. The suitability of a radiopharmaceutical depends, in part, upon the organ or tissue to be imaged. One particularly common choice is fluorodeoxyglucose (FDG), which is a molecule of glucose where a hydroxyl group is substituted with 18F. 18F is a β+ emitter meaning that it undergoes the following nuclear decay reaction:18F→18O+β++v+e−  eq. 1where β+ is a positron, v is a neutrino, and e− is an electron. The positron is ejected from the nucleus of 18F with substantial kinetic energy, which must be almost entirely dissipated before the positron can combine with an electron in an annihilation event. In general, dissipative processes can be elastic or inelastic scattering with any surrounding matter in the path of the positron, including electrons and nuclei. Statistically, positrons travel about 1 mm before losing enough kinetic energy to combine with an electron and annihilate. When annihilation occurs, a pair of 511 keV gamma photons is created equaling the energy equivalent of the annihilated particles, and radiating at close to 180° from each other. In the ideal case where the positron and electron both have zero momentum at the time of annihilation, the gamma photons would emit at exactly 180°. Deviation from 180° by about +/−0.5° indicates that the annihilation event occurred with particles having residual momentum.
It is known to place a pair of PET detectors 180° from each other to detect a pair of gamma photons emitted from a single annihilation event, and calculate the position of the annihilation event from the data collected. In some cases, the two or more PET detectors are rotated around the patient, and in others PET detectors form a continuous ring about the patient, thus requiring no rotation. In either case, the respective detectors collect gamma photons and either accept or reject the data depending in part on whether the photon was within an acceptable range about 511 keV, and whether it arrived within an acceptable time window to correlate one gamma photon to another. When a match is found between gamma photons, i.e. they are determined to have originated from the same annihilation event, a line of response (LOR) can be drawn between the two points on the respective detectors where the photons were detected. Accordingly, the position of the annihilation event must be located somewhere along the LOR.
Some instruments are capable of sufficient temporal resolution to calculate the position of the annihilation event based on the difference in the time of flight (TOF) of the pair of gamma photons. In lower resolution instruments other mathematical methods must be used to calculate annihilation position based on interpolation and/or extrapolation algorithms.
Traditionally PET/SPECT detectors include a plurality of scintillation crystals arranged in a pixelated two-dimensional array and spaced apart with septum material, which limits optical interference between adjacent crystals. The array of scintillation crystals is placed in optical communication with a plurality of photoconverters also arranged in a two-dimensional array. Often, one photoconverter will be in optical communication with a plurality of scintillation crystals. When a scintillation crystal receives a gamma photon, the photon travels some finite distance within the crystal before finally being absorbed. This distance, is known as the depth of interaction (DOI). At the position where the gamma photon is absorbed, the crystal emits a large number of UV and/or visible photons, i.e. it scintillates. The photon wave front propagates within the crystal and contacts the photoconverters. Traditionally, the photoconverters continuously integrate the photonic signal and are read individually based on whether they reach a minimum threshold signal intensity, and the data may be digitized thereafter. A center-of-mass calculation is then used to estimate the position of the scintillation event. From this data, parameters can be calculated for image reconstruction. For example, known image reconstruction algorithms can then be applied to the data to create an image. Such image reconstruction algorithms can include Filtered Back Projection and/or Ordered Subset Expectation Maximization. The reconstructed image can then be displayed according to known image display algorithms such as maximum intensity projection (MIP) and/or minimum intensity projection (mIP).
What is needed is an unpixelated slab detector, which would enable more precise calculations of, for example, depth of interaction and/or other parameters necessary to reconstruct high resolution images. Some embodiments of the present invention overcome one or more limitations of the prior art.